Electromagnetic shield for a passive electronic component in an active medical device implantable lead

ABSTRACT

A shielded component or network for an active medical device (AMD) implantable lead includes an implantable lead having a length extending from a proximal end to a distal end, all external of an AMD housing, a passive component or network disposed somewhere along the length of the implantable lead, the passive component or network including at least one inductive component having a first inductive value, and an electromagnetic shield substantially surrounding the inductive component or the passive network. The first inductive value of the inductive component is adjusted to account for a shift in its inductance to a second inductive value when shielded.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation application to U.S. application Ser.No. 13/860,888, filed on Apr. 11, 2013, which is continuationapplication of U.S. application Ser. No. 12/891,292, filed on Sep. 27,2010, now U.S. Pat. No. 8,437,865, which is a continuation-in-partapplication of U.S. application Ser. No. 12/873,862, filed on Sep. 1,2010, now U.S. Pat. No. 8,224,440, which is a continuation-in-partapplication of U.S. application Ser. No. 12/607,234, filed on Oct. 28,2009, now U.S. Pat. No. 8,175,700, which is a continuation-in-partapplication of U.S. application Ser. No. 12/407,402, filed on Mar. 19,2009, now U.S. Pat. No. 8,195,295, which is a continuation-in-partapplication of U.S. application Ser. No. 11/558,349, filed on Nov. 9,2006, now U.S. Pat. No. 7,945,322.

Furthermore, this application is a continuation-in-part application toU.S. application Ser. No. 12/788,123, filed on May 26, 2010, which is acontinuation-in-part application of U.S. application Ser. No.12/686,137, filed on Jan. 12, 2010, which is a continuation-in-partapplication to U.S. application Ser. No. 12/489,921, filed on Jun. 23,2009, now U.S. Pat. No. 7,751,903, which claims priority to U.S. Pro.App. Ser. Nos. 61/144,102, filed on Jan. 12, 2009 and 61/149,833, filedon Feb. 4, 2009.

FIELD OF THE INVENTION

This invention generally relates to the problem of high frequency energyinduced onto implanted leads during medical diagnostic procedures suchas magnetic resonant imaging (MRI). More specifically, the presentinvention relates to an implantable medical system comprised of anactive medical device (AMD) and at least one lead extending exteriorlyfrom a proximal end at or adjacent to the AMD, to a biological sensingor stimulating electrode at a distal end. The lead has a passivecomponent or network, including at least one inductive componentdisposed somewhere along its length between the proximal end and distalend. At least the inductive component of the passive component ornetwork is electromagnetically shielded.

BACKGROUND OF THE INVENTION

The radio frequency (RE) pulsed field of MRI can couple to an implantedlead in such a way that electromagnetic forces (EMFs) are induced in thelead. The amount of energy that is induced is related to a number ofcomplex factors, but in general, is dependent upon the local electricfield that is tangent to the lead and the integral of the electric fieldstrength along the lead. In certain situations, these EMFs can causecurrents to flow into distal electrodes or in the electrode interfacewith body tissue. It has been documented that when this current becomesexcessive, overheating of said lead or its associated electrode oroverheating of the associated interface with body tissue can occur.There have been cases of damage to such body tissue which has resultedin loss of capture of cardiac pacemaking pulses or tissue damage severeenough to result in brain damage or multiple amputations, and the like.

Electromagnetic interference (EMI) is also a significant issue. It hasbeen well demonstrated through various incidents and publications thatan implanted lead can act as an antenna and pick up unwanted signalsfrom the patient environment. In the past, there have been problems withmicrowave ovens, cell phones, and the like. Stray signals that arepicked up on implanted leads can be coupled to the interior of the AMDand interfere with sensitive electronic circuits. In cardiac pacemakers,instances of EMI being detected as normal cardiac rhythms have resultedin pacemaker inhibition which can be life-threatening.

Magnetic resonance imaging (MRI) is one of medicine's most valuablediagnostic tools. MRI is, of course, extensively used for imaging, butis also used for interventional medicine (surgery). In addition, MRI isused in real time to guide ablation catheters, neurostimulator tips,deep brain probes and the like. An absolute contra-indication forpacemaker or neurostimulator patients means that these patients areexcluded from MRI. This is particularly true of scans of the thorax andabdominal areas. Because of MRI's incredible value as a diagnostic toolfor imaging organs and other body tissues, many physicians simply takethe risk and go ahead and perform MRI on a pacemaker patient. Theliterature indicates a number of precautions that physicians should takein this case, including limiting the power of the MRI RF pulsed field(Specific Absorption Rate—SAR level), programming the pacemaker to fixedor asynchronous pacing mode, and then careful reprogramming andevaluation of the pacemaker and patient after the procedure is complete.There have been reports of latent problems with cardiac pacemakers orother AMDs after an MRI procedure, sometimes occurring many days later.Moreover, there are a number of papers that indicate that the SAR levelis not entirely predictive of the heating that would be found inimplanted leads or devices. For example, for magnetic resonance imagingdevices operating at the same magnetic field strength and also at thesame SAR level, considerable variations have been found relative toheating of implanted leads. It is speculated that SAR level alone is nota good predictor of whether or not an implanted device or its associatedlead system will overheat.

There are three types of electromagnetic fields used in an MRI unit. Thefirst type is the main static magnetic field designated B.sub.0 which isused to align protons in body tissue. The field strength varies from 0.5to 3.0 Tesla in most of the commonly available MRI units in clinicaluse. Some of the newer research MRI system fields can go as high as 11.7Tesla.

The second type of field produced by magnetic resonance imaging is thepulsed RF field which is generated by the body coil or head coil. Thisis used to change the energy state of the protons and elicit MRI signalsfrom tissue. The RF field is homogeneous in the central region and hastwo main components: (1) the electric field is circularly polarized inthe actual plane; and (2) the H field, sometimes generally referred toas the net magnetic field in matter, is related to the electric field byMaxwell's equations and is relatively uniform. In general, the RF fieldis switched on and off during measurements and usually has a frequencyof 21 MHz to 64 MHz to 128 MHz depending upon the static magnetic fieldstrength. The frequency of the RF pulse for hydrogen scans varies by theLamour equation with the field strength of the main static field where:RF PULSED FREQUENCY in MHz=(42.56) (STATIC FIELD STRENGTH IN TESLA).There are also phosphorous and other types of scanners wherein theLamour equation would be different.

The third type of electromagnetic field is the time-varying magneticgradient fields designated Bx, By, Bz, which are used for spatiallocalization. These change their strength along different orientationsand operating frequencies on the order of 1 kHz. The vectors of themagnetic field gradients in the X, Y and Z directions are produced bythree sets of orthogonally positioned coils and are switched on onlyduring the measurements.

At the frequencies of interest in MRI, RF energy can be absorbed andconverted to heat. The power deposited by RF pulses during MRI iscomplex and is dependent upon the power (Specific Absorption Rate (SAR)Level) and duration of the RF pulse, the transmitted frequency, thenumber of RF pulses applied per unit time, and the type of configurationof the RF transmitter coil used. The amount of heating also depends uponthe volume of tissue imaged, the electrical resistivity of tissue andthe configuration of the anatomical region imaged. There are also anumber of other variables that depend on the placement in the human bodyof the AMD and the length and trajectory of its associated lead(s). Forexample, it will make a difference how much EMF is induced into apacemaker lead system as to whether it is a left or right pectoralimplant. In addition, the routing of the lead and the lead length arealso very critical as to the amount of induced current and heating thatwould occur.

The cause of heating in an MRI environment is twofold: (a) RF fieldcoupling to the lead can occur which induces significant local heating;and (b) currents induced between the distal tip and tissue during MRI RFpulse transmission sequences can cause local Ohms Law heating in tissuenext to the distal tip electrode of the implanted lead. The RF field ofan MRI scanner can produce enough energy to induce RF voltages in animplanted lead and resulting currents sufficient to damage some of theadjacent myocardial tissue. Tissue ablation (destruction resulting inscars) has also been observed. The effects of this heating are notreadily detectable by monitoring during the MRI. Indications thatheating has occurred would include an increase in pacing capturethreshold (PCT), venous ablation, Larynx or esophageal ablation,myocardial perforation and lead penetration, or even arrhythmias causedby scar tissue. Such long term heating effects of MRI have not been wellstudied yet for all types of AMD lead geometries. There can also belocalized heating problems associated with various types of electrodesin addition to tip electrodes. This includes ring electrodes or padelectrodes. Ring electrodes are commonly used with a wide variety ofabandoned implanted device leads including cardiac pacemakers, andneurostimulators, and the like. Pad electrodes are very common inneurostimulator applications. For example, spinal cord stimulators ordeep brain stimulators can include a plurality of pad electrodes to makecontact with nerve tissue. A good example of this also occurs in acochlear implant. In a typical cochlear implant there would be sixteenpad electrodes placed up into the cochlea. Several of these padelectrodes make contact with auditory nerves.

Variations in the pacemaker lead length and implant trajectory cansignificantly affect how much heat is generated. A paper entitled,HEATING AROUND INTRAVASCULAR GUIDEWIRES BY RESONATING RF WAVES byKonings, et al., Journal of Magnetic Resonance Imaging, Issue 12:79-85(2000), does an excellent job of explaining how the RF fields from MRIscanners can couple into implanted leads. The paper includes both atheoretical approach and actual temperature measurements. In aworst-case, they measured temperature rises of up to 74 degrees C. after30 seconds of scanning exposure. The contents of this paper areincorporated herein by reference.

The effect of an MRI system on the leads of pacemakers, ICDs,neurostimulators and the like, depends on various factors, including thestrength of the static magnetic field, the pulse sequence, the strengthof RF field, the anatomic region being imaged, and many other factors.Further complicating this is the fact that each patient's condition andphysiology is different and each lead implant has a different lengthand/or implant trajectory in body tissues. Most experts still concludethat MRI for the pacemaker patient should not be considered safe.

It is well known that many of the undesirable effects in an implantedlead system from MRI and other medical diagnostic procedures are relatedto undesirable induced EMFs in the lead system and/or RF currents in itsdistal tip (or ring) electrodes. This can lead to overheating of bodytissue at or adjacent to the distal tip.

Distal tip electrodes can be unipolar, bipolar, multipolar and the like.It is very important that excessive RF current not flow at the interfacebetween the lead distal tip electrode or electrodes and body tissue. Ina typical cardiac pacemaker, for example, the distal tip electrode canbe passive or of a screw-in helix type as will be more fully described.In any event, it is very important that excessive RF current not flow atthis junction between the distal tip electrode and, for example, intosurrounding cardiac or nerve tissue. Excessive current at the distalelectrode to tissue interface can cause excessive heating to the pointwhere tissue ablation or even perforation can occur. This can belife-threatening for cardiac patients. For neurostimulator patients,such as deep brain stimulator patients, thermal injury can causepermanent disability or even be life threatening. Similar issues existfor spinal cord stimulator patients, cochlear implant patients and thelike.

A very important and possibly life-saving solution is to be able tocontrol overheating of implanted leads during an MRI procedure. A noveland very effective approach to this is to first install parallelresonant inductor and capacitor bandstop filters at or near the distalelectrode of implanted leads. For cardiac pacemaker, these are typicallyknown as the tip and ring electrodes. One is referred to U.S. Pat. No.7,363,090; US 2007/0112398 A1; US 2008/0071313 A1; US 2008/0049376 A1;US 2008/0024912 A1; US 2008/0132987 A1; and US 2008/0116997 A1, thecontents of all of which are incorporated herein. US 2007/0112398 A1relates generally to L-C bandstop filter assemblies, particularly of thetype used in active implantable medical devices (AIMDs) such as cardiacpacemakers, cardioverter defibrillators, neurostimulators and the like,which raise the impedance of internal electronic or related wiringcomponents of the medical device at selected frequencies in order toreduce or eliminate currents induced from undesirable electromagneticinterference (EMI) signals.

Other types of component networks may also be used in implantable leadsto raise their impedance at MRI frequencies. For example, a seriesinductor may be used as a single element low pass filter. The inductancewill tend to look like a high impedance at high frequencies, such as theRF pulsed frequencies of a typical MRI scanner. For more information onthis refer to U.S. Pat. No. 5,217,010 (Tsitlik et al.), the contents ofwhich are incorporated herein by reference.

U.S. Pat. No. 7,363,090 and US 2007/0112398 A1 show resonant L-Cbandstop filters placed at the distal tip and/or at various locationsalong the medical device leads or circuits. These L-C bandstop filtersinhibit or prevent current from circulating at selected frequencies ofthe medical therapeutic device. For example, for an MRI system operatingat 1.5 Tesla, the pulsed RF frequency is 63.84 MHz, as described by theLamour Equation for hydrogen. The L-C bandstop filter can be designed toresonate at or near 63.84 MHz and thus create a high impedance (ideallyan open circuit) in the lead system at that selected frequency. Forexample, the L-C bandstop filter when placed at the distal tip electrodeof a pacemaker lead will significantly reduce RF currents from flowingthrough the distal tip electrode and into body tissue. The L-C bandstopfilter also reduces EMI from flowing in the leads of a pacemaker therebyproviding added EMI protection to sensitive electronic circuits. Ingeneral, the problem associated with implanted leads is minimized whenthere is a bandstop filter placed at or adjacent to its distal tipelectrodes.

At high RF frequencies, an implanted lead acts very much as like anantenna and a transmission line. An inductance element disposed in thelead will change its transmission line characteristics. The inductancecan act as its own antenna pick-up mechanism in the lead and therefore,ideally, should be shielded. When one creates a very high impedance atthe distal electrode to tissue interface by installation of a resonantbandstop filter as described in U.S. Pat. No. 7,038,900 and as furtherdescribed in US 2007/0112398 A1, there is created an almost open circuitwhich is the equivalent of an unterminated transmission line. Thiscauses a reflection of MRI induced RE energy back towards the proximalend where the AIMD (for example, a pacemaker) is connected. In order tocompletely control the induced energy in an implanted lead, one musttake a system approach. In particular, a methodology is needed wherebyenergy can be dissipated from the lead system at the proximal end in away that does not cause overheating either at the distal electrodeinterface or at the proximal end cap. Maximizing energy transfer from animplanted lead is more thoroughly described in US 2010/0160997 A1, thecontents of which are incorporated herein by reference.

Accordingly, there is a need for attenuating the RF energy that can beinduced onto or into an implanted lead system.

Further, there is a need to provide shielding of passive networkcomponents, including any inductors that would be disposed along thelength of the lead. Such shielding should reduce or prevent externalelectromagnetic fields from coupling RF electromagnetic energy to saidpassive component or network and, in particular, its inductivecomponent(s). The present invention fulfills these needs and providesother related advantages.

SUMMARY OF THE INVENTION

The present invention resides in a shielded component or network for anactive medical device (AMD) implantable lead, comprising: (1) animplantable lead having a length extending from a proximal end to adistal end, all external of an AMD housing, (2) a passive component ornetwork disposed somewhere along the length of implantable lead, thepassive component or network including at least one inductive component,having a first inductive value, and (3) an electromagnetic shieldsubstantially surrounding the inductive component or the passivenetwork. The first inductive value of the inductive component isadjusted to account for a shift in its inductance to a second inductivevalue when shielded.

The passive component network may include at least one capacitivecomponent electrically connected in parallel with the at least oneinductive component to form a bandstop filter. The inductive componentmay comprise a solenoid inductor or a chip inductor, and the capacitivecomponent may comprise a chip capacitor, parasitic capacitance, or afeedthrough capacitor. In this regard, the capacitive component maycomprise parasitic capacitance formed between coils of the inductivecomponent and/or between the inductive component and the electromagneticshield. In a preferred embodiment, a dielectric material is disposedbetween coils of the inductive component and between the inductivecomponent and the electromagnetic shield to facilitate formation of theparasitic capacitance. The capacitive component and the inductivecomponent may form a parallel resonant bandstop filter and are tuned toimpede induced current flow through the implantable lead at a selectedcenter frequency or range of frequencies, typically comprising an MRI RFpulsed frequency or range of RF pulsed frequencies. The MRI RF pulsedfrequency range includes tens of kilobertz, hundreds of kilohertz ormegahertz.

A non-conductive insulator or dielectric material may be disposedbetween the passive component network and the electromagnetic shield.

In a preferred embodiment, an inductor is provided having first andsecond conductive terminals in spaced non-conductive relation, and acapacitor is also provided having first and second conductive terminalsin spaced non-conductive relation. The inductor and the capacitor arephysically disposed in series relative to one another, and areelectrically connected to one another in parallel to form a bandstopfilter. One of the first or second conductive terminals of the inductoris disposed generally adjacent to one of the first or second conductiveterminals of the capacitor. The capacitor and the inductor may bealigned along a common axis, and the adjacent conductive terminals ofthe inductor and the capacitor may abut one another. An electricalinsulator may also be disposed between the adjacent conductive terminalsof the inductor and the capacitor.

The electrical potential between the adjacent conductive terminals ofthe inductor and the capacitor is preferably minimized, and is, ideally,zero.

The second conductive terminal of the inductor may be conductivelycoupled to the first conductive terminal of the capacitor, and the firstconductive terminal of the inductor may be conductively coupled to thesecond conductive terminal of the capacitor.

A plurality of paired inductor and capacitor bandstop filters may beprovided, wherein each bandstop filter is physically disposed in seriesrelative to one another. In this case, each paired inductor andcapacitor is electrically connected in series to another paired inductorand capacitor.

The capacitive component and the inductive component may comprisebiocompatible and non-migratable materials and/or they may be disposedwithin a medically sealed container which forms the electromagneticshield. The hermetically sealed container also forms an electromagneticshield and may comprise a biocompatible housing in which the bandstopfilter is disposed, and biocompatible first and second conductivecontacts extending through and in non-conductive relation with thehousing, which are conductively coupled in series to the bandstopfilter. Typically, the hermetically sealed container is disposed inseries in the implantable lead, wherein first and second contacts of thehermetically sealed container are connected to, respectively, proximaland distal portions of the lead.

A substrate may be provided onto which the inductor and capacitor arefixed in a pre-assembly prior to insertion into the biocompatible shieldhousing. First and second hermetic terminals hermetically sealed to thebiocompatible housing after the pre-assembly is inserted therein maycomprise at least a portion of the first and second conductive contacts,respectively. An electrically insulated conformal coating may be appliedover at least a portion of the hermetically sealed container.

The overall Q of the band stop filter is selected to balance impedanceat the selected frequency versus frequency bandwidth characteristics.When the Q of the inductive component is relatively high, the Q of thecapacitive component is relatively low such that the inductive componenthas a relatively low resistive loss and the capacitive component has arelatively high equivalent series resistance. When the Q of theinductive component is relatively low and the Q of the capacitivecomponent is relatively high, the inductive component has a relativelyhigh resistive loss and the capacitive component has a relatively lowequivalent series resistance.

The active medical device (AMD) has a conductive equipotential surface,wherein the electromagnetic shield that substantially surrounds theinductive element or the passive network is conductively coupled toeither the AMD equipotential surface or to surrounding body tissue. TheAMD equipotential surface may comprise a conductive biocompatiblehousing for the AMD.

An energy diversion circuit may be provided which conductively couplesthe implantable lead to the electromagnetic shield. The energy diversioncircuit may comprise a low pass filter such as a capacitor, an inductor,a Pi filter, a T filter, an LL filter or an “n” element filter. Theenergy diversion circuit may further comprise at least one seriesresonant L-C trap filter.

The energy diversion circuit may also comprise a high pass filter whichprevents low frequency radiant field-induced energy in the implantedlead from passing through the diversion circuit to an energy dissipatingsurface or ground. The high pass filter may comprise a capacitor, aresistor in series with a capacitor, or an L-C trap filter.

An impeding circuit may be provided for raising the high frequencyimpedance of the implantable lead. The impedance circuit may comprise aninductor and/or a bandstop filter.

The electromagnetic shield may comprise the energy dissipating surface.

The inductive component may comprise a plurality of spaced apartinductive components disposed along the length of the implantable lead.In this case, not all of the inductive components need be shielded.Further, the electromagnetic shield may comprise a plurality ofelectromagnetic shields disposed along the length of the implantablelead. An adjacent pair of the plurality of electromagnetic shields aretypically spread apart from one another but are also typicallyconductively coupled to one another.

The electromagnetic shield may comprise a conductive heat-shrink tubing,a conductive foil, wire, braid, mesh, circuit trace, or solid tubularmaterial, or a conductive polymer, a conductive epoxy, carbonnano-fibers, nano-meshes, nano-coatings or nano-threads. Theelectromagnetic shield is further typically radially spaced from thepassive component network.

The electromagnetic shield may comprise MP35N, iridium, carbon,platinum, titanium, palladium, chromium, Wolfram, tungsten, gold,copper, or alloys thereof. The inductive component may comprise a chipinductor, a solenoid inductor, a Wheeler spiral or a circuit traceinductor. Moreover, the implantable lead may comprise a plurality ofimplantable leads substantially surrounded by the electromagneticshield. The AMD may comprise an implantable hearing device, a cochlearimplant, a pisoelectric soundbridge transducer, a neurostimulator, abrain stimulator, a cardiac pacemaker, a left ventricular assist device,an artificial heart, a drug pump, an implantable bone growth stimulator,a urinary incontinence device, a spinal cord stimulator, an anti-tremorstimulator, an implantable cardioverter defibrillator, or a congestiveheart failure device.

Empirical data may be used to adjust the value of the inductance fromthe first inductive value to the second inductive value. Alternatively,an equivalent circuit model, such as PSPICE may be utilized to adjustthe value of the inductance from the first inductive value to the secondinductive value. Further, mathematical formulations based on amagnetostatic integral equation may be utilized to adjust the value ofthe inductance from the first inductive value to the second inductivevalue. Moreover, this ratio may be applied over the empirical inductanceand divided by the inductance in the electromagnetic shield to adjustthe value of the inductance from the first inductive value to the secondinductive value.

The electromagnetic shield may comprise a housing for a passive fixationtip electrode or may be associated with a translational active fixationtip, wherein the housing for the active fixation tip comprises theelectromagnetic shield. Alternatively, the electromagnetic shield may bedisposed within a housing for the active fixation tip.

The network may include an active electronic circuit.

The shield may comprise a non-metallic material such as sapphire, ruby,alumina and/or ceramic materials which have a thin conductive coatingdeposited thereon by plating, chemical vapor deposition, sputtering,physical application, cladding or the like.

Other features and advantages of the present invention will becomeapparent from the following more detailed description, taken inconjunction with the accompanying drawings which illustrate, by way ofexample, the principles of the invention.

BRIEF DESCRIPTION OF THE DRAWINGS

The accompanying drawings illustrate the invention. In such drawings:

FIG. 1 is a wire-formed diagram of a generic human body showing a numberof exemplary active medical devices (AMDs);

FIG. 2 illustrates an exemplary prior art cardiac pacemaker with theleads schematically shown extending to a patient's heart;

FIG. 3 is a schematic illustration of a prior art AMD with a bipolarlead;

FIG. 4 is similar to FIG. 3, except that the bipolar lead wires arecoaxially wound around one another;

FIG. 5 is an enlarged schematic view of the area indicated by line 5-5from FIG. 4, and illustrates an inductor disposed in series with each ofa pacemaker tip and ring electrode circuits;

FIG. 6 is similar to FIG. 5 except that the inductor elements have beenreplaced by bandstop filters;

FIG. 7 is taken of the area indicated by line 7-7 from FIG. 4 and issimilar to FIGS. 5 and 6, except that an overall shield encompassesvarious impeder and diverter elements;

FIG. 8 is similar to FIG. 7, and illustrates a diverter elementconsisting of an inductor in series with a capacitor to form an L-C trapfilter;

FIG. 9 illustrates a probe or catheter which has a shielded sectionembodying the present invention;

FIG. 10 is an enlarged sectional view taken along line 10-10 from FIG.9;

FIG. 11 is an enlarged view of the distal tip section of the probe orcatheter of FIG. 9;

FIG. 12 is similar to FIG. 5 except that it has a single shield for adifferential mode L-C trap filter;

FIG. 13 is similar to FIG. 12, except that the L-C trap filter has beenreplaced by a general diverter element;

FIG. 14 is similar to FIG. 13, wherein the impeder elements areinductors;

FIG. 15 is similar to FIGS. 1.3 and 14, wherein the impeder elements arebandstop filters;

FIG. 16 is a graph showing impedance versus frequency for the idealbandstop filter circuit of FIGS. 6 and 15;

FIG. 17 is a schematic illustration similar to FIG. 3, showing a genericunipolar AMD and lead with a bandstop filter added at or near a distalelectrode;

FIG. 18 illustrates a family of curves which show the attenuation in dBversus frequency for bandstop filters;

FIG. 19 illustrates a hermetically sealed container for a bandstopfilter embodying the present invention;

FIG. 20 is an enlarged sectional view taken generally along line 20-20from FIG. 19;

FIG. 21 is an enlarged perspective view of the hermetic seal assemblyfrom FIG. 20;

FIG. 22 is an enlarged cross-sectional view taken generally along theline 22-22 from FIG. 20;

FIG. 23 is an end view of the hermetic seal assembly of FIG. 21, takenfrom line 23-23 in FIG. 20;

FIG. 24 is a perspective view illustrating a multi-layer flex cable ontowhich the inductor and capacitor of FIG. 20 are mounted;

FIG. 25 is a schematic illustration showing that the inductor andcapacitor are physically disposed in series relative to one another andyet electrically connected in parallel;

FIG. 26 is an electrical schematic diagram of the bandstop filter ofFIGS. 24 and 25;

FIG. 27 is a sectional view similar to FIG. 20, but illustrating analternative embodiment where the chip capacitor has been replaced with afeedthrough capacitor;

FIG. 28 is a schematic illustration showing electrical connections ofthe inductor and capacitor relative to the lead;

FIG. 29 is an electrical schematic diagram of the structure shown inFIGS. 27 and 28;

FIG. 30 is a perspective view of a solenoid inductor wrapped around anon-ferromagnetic core;

FIG. 31 is similar to FIG. 30, showing inductor wires coiled around aplastic support structure to form the equivalent of an air-woundinductor;

FIG. 32 is a sectional view taken along the line 32-32 from FIG. 30,illustrating the magnetic field when no shield is present;

FIG. 33 is similar to FIG. 32, illustrating changes in the magneticfield when the inductor is shielded;

FIG. 34 illustrates a thick film inductor;

FIG. 35 shows the arrangement of circuit traces that form the thick filminductor of FIG. 34;

FIG. 36 is similar to FIG. 32, illustrating electromagnetic field linesfor the thick film inductor when not shielded;

FIG. 37 is similar to FIG. 33, illustrating changes in the magneticfield lines when the thick film inductor is shielded;

FIG. 38 illustrates distribution of current induction for a thick filminductor when shielded;

FIG. 39 is a graph of histograms from various prototypes of solenoidinductors;

FIG. 40 is a graph of impedance versus frequency curves for circuitsboards with either thick film chip or wire wound or solenoid inductors,when shielded in comparison to when not shielded;

FIG. 41 is a PSPICE computer model for predicting the resonant circuitbehavior of a solenoid inductor in air;

FIG. 42 is a PSPICE computer model similar to FIG. 41, modified to addmutual inductance coupling to a surrounding electromagnetic shield;

FIG. 43 is a graph of the frequency response predicted by the PSPICEmodels of FIGS. 41 and 42;

FIG. 44 is a perspective view of a passive electrode fixation tiptypically used in cardiac pacemaker applications;

FIG. 45 is an enlarged sectional view taken generally along line 45-45from FIG. 44;

FIG. 46 is an electrical schematic diagram for the circuit of FIG. 45;

FIG. 47 is a perspective view of a typical off-the-shelf commercialmonolithic ceramic capacitor (MLCC);

FIG. 48 is a perspective view of a typical off-the-shelf commercialunipolar feedthrough capacitor;

FIG. 49 is a sectional view similar to FIG. 45, except that the inductorelement is wire wound around a non-ferromagnetic mandrel;

FIG. 50 is a sectional view similar to FIGS. 45 and 49 wherein a pair ofinductors are coupled in connection with a capacitor to form a “T”filter within the passive electrode tip;

FIG. 51 an electrical schematic for the structure shown in FIG. 50;

FIG. 52 is a perspective view of a generic prior art active fixationdistal tip typically used in conjunction with cardiac pacemakers;

FIG. 53 is an enlarged sectional view taken generally along line 53-53from FIG. 52;

FIG. 54 is a fragmented sectional view of a portion of the activefixation tip of FIGS. 52 and 53, modified to include a shielded inductoror bandstop filter in accordance with the present invention;

FIG. 55 is a sectional view similar to FIG. 54, where an inductive coilis disposed within a dielectric material such that parasiticcapacitances form, with the inductor coil itself, a bandstop filter;

FIG. 56 is an equivalent cross-section schematic diagram for thestructure shown in FIG. 55;

FIG. 57 is an electrical schematic for the structure shown in FIGS. 55and 56;

FIG. 58 is a simplified electrical schematic of the structure shown inFIGS. 55-57;

FIG. 59 is a perspective view of a reinforced polyimide tubing thatincludes the shielding for an inductive component in accordance with thepresent invention;

FIG. 60 is an enlarged sectional view taken generally along 60-60 fromFIG. 59;

FIG. 61 is similar to FIG. 59, illustrating an alternative embodimentwhere an insulation tube is slipped over the lead and then a shieldlayer is slipped over the insulation tube;

FIG. 62 is similar to FIG. 61, except that the metal shield tube isreplaced by wire wound strands;

FIG. 63 is similar to FIGS. 61 and 62, except that the metal shield tubeor wire wound strands are replaced by wrapped foil;

FIG. 64 shows an open mesh cross-braided shield wire instead of thewound shield wire of previous embodiments; and

FIG. 65 is an enlarged perspective view of the cross-braided shield ofFIG. 64.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

As shown in the drawings for purposes of illustration, the presentinvention relates to a system for RF shielding of a passive component ornetwork disposed along the length of implanted leads of active medicaldevices (AMDs). In particular, the RF shielding is to shield a passiveinductor or inductive component in the presence of high powerelectromagnetic field environments, such as the RF pulsed fieldsproduced by a clinical MRI scanner. In a broad sense, the presentinvention comprises an active medical device system including animplanted lead having RF shielded inductors or passive networkcomponents including inductors. The implanted lead may be coaxial,rectangular, flat or other geometries. US 2009/0243756 incorporatedherein by reference. Furthermore, the implanted lead may consist of anumber of internal conductors, such as a bipolar lead for cardiacpacemaker channel or even an eight or sixteen conductor spinal cordstimulator implanted lead. This is also known as a multichannel leadsystem. The networks of the present invention can also include activecomponents in combination with at least one inductor. These activenetworks can comprise a portion of a lead-based sensor, such as ahemodynamic sensor, a pulse oxygen sensor, an acceleration or ratesensor, and the like. In addition, the inductors of the presentinvention can be a part of an energy harvesting device or circuit whichis implanted in the human body. US 2009/0243756 is incorporated byreference herein.

In general, the shield of the present invention surrounds all of thepassive or active component elements disposed along the length of animplanted lead, including, but not limited to, inductors,inductor-capacitor (L-C) bandstop filters, L-C trap filters, or singleor multi-element low pass filters. It is important that the shieldsurround at least the inductor component(s) associated with suchelectronic networks. As a practical matter, the shield would generallyencompass all of the passive components. The shield could also surroundall of the conductors in a particular implanted lead that is routed to aparticular area of body tissue. For example, in a cardiac pacemakerapplication, there are often dual chamber bipolar conductors in theimplanted lead, wherein one lead is typically routed to the rightventricle and the other to the right atrium. Each of these implantedleads, consisting of two conductors, would have its own passivecomponent filtering elements which would be individually shielded.Typically, conforming to the shape of the leads, the shields of thepresent invention may be coaxial, flat, rectangular or any othergeometry suitable for either tunneling or for transvenous insertionwithin the human body.

The shield of the present invention can also act as an energydissipating surface. Diverting circuits, consisting of eithercapacitors, low-pass filter, L-C trap filters or high-pass filters, canbe used to divert energy from an implanted lead to its surroundingshield. The shield, in a preferred embodiment, is in contact with bodytissue whereby induced RF energy from the lead is diverted to theshield, which in turn acts as an energy dissipating surface (EDS). US2010/002300 A1 is incorporated herein by reference.

Implanted leads have both a characteristic impedance and also act as atransmission line. They tend to effectively couple energy from anexternal electromagnetic interference emitter as a function of theirwavelength. This also varies with lead trajectory, design and otherfactors. However, when one is only concerned with particular frequencyranges, for example the RF pulse frequency of MRT, it is not necessaryto shield the entire lead. In this regard, one could shield asignificant portion of the lead so that the exposed (unshielded) portionof the lead was significantly less than a half or a quarter wavelengthin body tissue. This makes the remaining unshielded lead portion a veryinefficient antenna and therefore it would only pick up a very smallamount of induced energy. Accordingly, in accordance with the presentinvention, one could shield passive network components or inductancesdisposed along the length of the shield and could also shield adjoiningsections of the lead itself. By shielding a portion of the implantedlead or even segments of the implanted lead, one would break up itsresonant lengths thereby making it a very ineffective antenna over abroad range of MRI pulsed frequencies.

The shields of the present invention can be a solid conductor, woundspiral conductors, meshes, tubing, nano-coatings or the like. In thepreferred embodiment, the shield would present a fairly homogenousconductive surface such that it would effectively reflect and/or absorbincident electromagnetic fields. However, complete shielding is reallynot necessary. Accordingly, the shield could be loosely woven such thatonly a portion of the electromagnetic interference was intercepted.

The invention further resides in a combination of shields with one ormore impeding circuits which could also be optimally combined with oneor more diversion circuits. The impeding circuits typically wouldconsist of either inductors or L-C parallel resonant-bandstop filters.The diversion circuits would typically consist of a capacitor, amulti-element low-pass filter, a high-pass filter, or an L-C trapfilter. The operation of impeding circuits and diversion circuits ismore thoroughly described in US 2010/002300 A1 and US 2010/0160997 A1,which are incorporated by reference. In a particularly preferredembodiment, the shield of the present invention is used in combinationwith an impeding circuit known as a bandstop filter. The bandstop filterhas a Q and 3-dB bandwidth such that, at resonance, it offersattenuation of at least 10 dB over a range of MRI RF pulsed frequenciesat least 100 kHz wide.

In the case where bandstop filters are installed at or near the distalelectrode of an implanted lead, the RF energy induced by the MRI pulsefield is inhibited from flowing into body tissues and thereby beingdissipated. However, even when distal electrode bandstop filters areused, that energy still resides in the lead system. In other words, bypreventing this induced energy from flowing to sensitive tissues atdistal electrode interfaces, a great deal has been accomplished;however, it is still important to carefully dissipate the remainingenergy that is trapped in the lead system.

In order to provide optimal decoupling of RF energy from an implantedlead to the energy dissipating surface of a shield, one should considerThevenin's maximum power transfer theorem. It is well known inelectrical engineering that to transfer maximum power to a load, theload impedance must be equal to the source impedance. If the sourceimpedance is completely resistive, for example, 50 ohms, then totransfer maximum power the load impedance would have to be 50 ohms. Whenthe source impedance is reactive, then to transfer maximum power toanother location the load impedance should have the opposite sign ofreactance and the same impedance and resistance. In a typical implantedlead system, the implanted leads typically appear inductive.Accordingly, having a capacitive energy diversion circuit to coupleenergy from the lead conductors to the EDS shield surface, one has atleast some cancellation of these imaginary impedance factors. Inelectrical engineering, the inductance of the lead would be denoted by+jΩL. The impedance of the capacitor, on the other hand, is a −j/ΩCterm. Transferring maximal energy from a lead to an energy dissipatingsurface is more thoroughly described in U.S. Pat. No. 7,689,288 thecontents of which are incorporated herein. It is important that theinductive elements of the diverter and/or impeding circuits be shieldedin accordance with the present invention.

FIG. 1 is a wire formed diagram of a generic human body showing a numberof exemplary implanted medical devices. 100A is a family of implantablehearing devices which can include the group of cochlear implants,piezoelectric sound bridge transducers and the like. 100B includes anentire variety of neurostimulators and brain stimulators. 100C shows acardiac pacemaker. 1000 includes the family of left ventricular assistdevices (LVAD's) and artificial hearts. 100E includes an entire familyof drug pumps which can be used for dispensing of insulin, chemotherapydrugs, pain medications and the like. 100F includes a variety ofimplantable bone growth stimulators for rapid healing of fractures. 100Gincludes urinary incontinence devices. 100H includes the family of painrelief spinal cord stimulators and anti-tremor stimulators. 100Iincludes a family of implantable cardioverter defibrillator (ICD)devices, congestive heart failure devices (CHF), and cardioresynchronization therapy devices, otherwise known as CRT devices. 110Jillustrates a family of probes or catheters that can be transvenouslyinserted during catheter lab procedures. These are normally consideredshort-term implants in that they are inserted within the human body forat most a few hours.

The various types of active medical devices (AMDs) illustrated in FIG. 1generally represent any type of AIMD that is considered a “long-term”implant. This is in direct contrast to things like probes or cathetersor surgical devices that are “short-term” body insertions. For example,a probe or catheter is typically used in a cath-lab situation wherein itis temporarily inserted through a femoral (or other) artery where theentire procedure lasts minutes or at most a few hours. On the otherhand, a long-term implant, such as a cardiac pacemaker, is generallydesigned to be implanted in the human body for many years. There aresignificant differences in the art between a short-term and a long-termimplant. For example, for a long-term implant, one has to worry greatlyabout the long-term biocompatibility, toxicity and even the hermeticityof the implant. In contrast, a probe, catheter or temporary looprecorder need only operate or be reliable for a matter of minutes oreven hours. In general, a short-term implant is often considered to be adisposable device. In addition, the FDA regulatory approval processesfor long-term implants is significantly different and involves much morerigorous testing and product safety and reliability criteria. The FDACenter for Devices and Radiological Health (FDA-CDRH) is the responsibleregulatory agency for long-term cardiac implants. As used herein, theterm active medical device (AMD) is construed to include long-termimplants and also short-term body insertions, such as probes orcatheters. The term AMD is inclusive of active implantable medicaldevices (AIMDs) and also externally worn medical devices that areassociated with an implanted lead.

Throughout, the term lead generally refers to implantable leads andtheir conductors that are external to the housing of the active medicaldevice. These leads tend to have a proximal end, which is at or adjacentto the AMD, and a distal end, which typically includes one or moreelectrodes which are in contact with body tissue.

FIG. 2 is a drawing of a typical cardiac pacemaker 1000 showing atitanium case or housing 102 and an IS-1 header connector block 104. Thetitanium case or housing 102 is hermetically sealed, however there is apoint where leadwires 106 a-106 d must ingress and egress a hermeticseal. This is accomplished by providing a hermetic terminal assembly 108that generally consists of a ferrule 110 which is laser welded to thetitanium housing 102 of the pacemaker 100C.

Four leadwires are shown consisting of leadwire pair 106 a and 106 b andleadwire pair 106 c and 106 d. This is typical of what is known as adual chamber bipolar cardiac pacemaker. The IS-1 connectors 112 and 112′of leads 114 and 114′ are designed to plug into receptacles 116 and 116in the header block 104. The receptacles 116 and 116′ are low voltage(pacemaker) connectors covered by an ANSI/AAMI ISO standard IS-1. Highervoltage devices, such as implantable cardioverter defibrillators (ICDs),are covered by ANSI/AAMI ISO standard DF-1. A new standard which willintegrate both high voltage and low voltage connectors into a miniaturein-line quadripolar connector is known as the IS-4 series. The implantedleads 114 and 114′ are typically routed transvenously in a pacemakerapplication down into the right atrium 118 and the right ventricle 118′of the heart 120. New generation biventricular or CRT-P devices mayintroduce leads to the outside of the left ventricle, which devices haveproven to be very effective in cardiac resynchronization and treatingcongestive heart failure (CHF).

Although the present invention will be described herein in the contextand environment of a cardiac pacemaker 100C and its associated leads 114and 114′, the present invention may also be advantageously utilized inmany other types of AMDs as briefly outlined above, as well as in othercommercial electronic, military, aerospace and other applications. Inthe following discussion, to the extent practicable, functionallyequivalent components will retain the same or a similar referencenumber, irrespective of the particular embodiment being described.

FIG. 3 illustrates a prior art single chamber bipolar device 100C andlead system 114 and 114′ with a distal tip electrode 122 and a ringelectrode 124 typically as used with the cardiac pacemaker 1000. Shouldthe patient be exposed to the fields of an MRT scanner or other powerfulemitter used during a medical diagnostic procedure, currents that aredirectly induced in the lead system 114, 114′ can cause heating by1.sup.2R losses in the lead system or by heating caused by RF currentflowing from the tip and ring electrodes 122, 124 into body tissue. Ifthese induced RF currents become excessive, the associated heating cancause damage or even destructive ablation to body tissue.

FIG. 4 illustrates a single chamber bipolar cardiac pacemaker 1000, andleads 114 and 114′ having distal tip 122 and distal ring 124 electrodes.This is a spiral wound (coaxial) system where the ring coil 114′ iswrapped around the tip coil 114. There are other types of pacemakerleadwire systems in which these two leads lay parallel to one another(known as a bifilar lead system), which are not shown.

FIG. 5 is taken from section 5-5 of FIG. 4 and illustrates an inductor Ldisposed in series with each of a pacemaker tip and ring electrodecircuits. The inductor L acts as a single element low pass filter andtends to attenuate the flow of current at high frequencies, such as MRIRF pulsed frequencies. The operation of inductors disposed inimplantable lead wires is more thoroughly described in U.S. Pat. No.5,217,010, the contents of which are incorporated herein. The shadedareas 126 in FIG. 5 illustrate that each of the inductors L.sub.1 andL.sub.2 has been shielded in accordance with the present invention. Thisprotects the inductors L.sub.1 and L.sub.2 from picking up strayelectromagnetic interference (EMI) from powerful RF fields of medicaldiagnostic equipment, such as an MRI scanner. For simplicity, theshields 126 are shown with a ground symbol (GND) indicating that theycan be grounded in a number of ways. The important thing is that theshield 126 be able to reflect, absorb and dissipate RF energy thatcouples onto it. Electromagnetic shields both absorb and reflectincident high frequency energy. The energy that is reflected is notcoupled onto the implanted lead. However, the energy that is absorbed isbest converted to heat and dissipated into surrounding body tissues.Another method of grounding the shields is to connect a conductor backto the conductive housing 102 of the AMD itself. In general, thearrangement for a cardiac pacemaker shown in FIG. 5 is preferable inthat the tip electrode inductor L.sub.2 and the ring electrode inductorL.sub.1 are individually shielded. It would be undesirable to have anoverall electromagnetic shield to shield both the tip inductor L.sub.2,the ring electrode 124 and the ring inductor L.sub.1. This is becauseimportant cardiac pacing and biological signal sensing functions occurbetween the electrodes 122 and 124. Overall shielding of both tip andring electrodes would impair said functions.

FIG. 6 is taken from section 6-6 of FIG. 4 and is very similar to FIG.5. In this case, the inductor elements L.sub.1 and L.sub.2 have beenreplaced by L-C bandstop filters 128 and 130. In general, the shieldedbandstop filters would be tuned to be resonant at a center frequency ina range of MRI RF pulsed frequencies. The operation of bandstop filtersin implanted leads is more thoroughly described by U.S. Pat. No.7,363,090 and US 2007/0112398, the contents of which are incorporatedherein by reference.

FIG. 7 illustrates an AMD bipolar lead system similar to that describedin FIGS. 5 and 6 except that an overall shield 126 encompasses variousimpeder elements 132 and 134 as well as diverter elements 136, 138 and140. In this case, the impeder elements 132 and 134 could be inductorsor bandstop filters as previously taught in FIGS. 5 and 6. The diverterelements could be capacitors or L-C trap filters as taught in U.S. Pat.No. 7,689,288, the contents of which are incorporated herein. In thiscase, both the tip electrode 122 and the ring electrode 124 are disposedoutside the single shield 126 so they can still perform their vitalcardiac pacing and biological sensing functions

FIG. 8 is very similar to FIG. 7 and illustrates a diverter element 140consisting of an inductor L in series with a capacitor C, forming whatis known as an L-C trap filter. The operation of these trap filters isdescribed in U.S. Pat. No. 7,689,288. The shield 126 protects theinductive component L of the L-C trap filter from picking up unwantedelectromagnetic interference, for example, in an MRI RF fieldenvironment.

FIG. 9 illustrates a probe or a catheter 100J which has a shieldedsection 142 which encompasses inductors or electronic networks of thepresent invention. The probe or catheter 100J consists of the flexibleand steerable probe or catheter section 144 which may be bent as shownand generally terminates in one or more distal electrodes 146. Thesedistal electrodes consist of mapping electrodes, ablation electrodes andthe like. There is generally a catheter handle or body 148 which is usedfor steering the probe or catheter into the body transvenously. Thesehandles can take the form of a pistol grip or many other shapes.

FIG. 10 is taken from section 10-10 of FIG. 9 and shows two leads insidethe flexible portion 144 of the probe or catheter. Shown are diverterelements 138 and 140 which are connected between each of the leads 114and 114′ to the electromagnetic shield 126 of the present invention. Ina preferred embodiment, the diverter elements would be L-C trap filterswhich would be used to divert energy picked up on the leads 114 and 114′to the shield surface 126. An optional insulation sleeve 150 is shown,which is generally undesirable. In other words, it is preferable thatthe conductive shield 126 be in contact with body tissue so that itdissipates unwanted MRI RF energy over a large surface area. Theelectromagnetic shield 126 of the present invention protects theinductor element of the L-C trap filters 138 and 140 from picking upunwanted high frequency RF energy.

In the description of the various embodiments shown in the accompanyingdrawings, the functionally equivalent components shall have the samereference number.

FIG. 11 is an alternative to the shielded portion 142 of the probe orcatheter 100J illustrated in FIGS. 9 and 10. In this case, there arethree internal conductors 114, 114′ and 114″ disposed within theflexible catheter portion 142. There are also three segmented shields126, 126′ and 126″ for a respective bandstop filter 128, 1283 and 128″.Each of the bandstop filters 128, 128′ and 128″ are connected in serieswith a respective one of the catheter conductors. In this case, theshields 126, 126′ and 126″ could be continuous or segmented as shown.There is an advantage to segmented shields as this promotes theflexibility of an implanted probe, catheter or AMD lead. In addition,segmented shield sections break up transmission line type resonances andchange the wavelength of the implanted lead to make it a much lessefficient RF antenna. One can see that there are sensing electrodes 122,122′ and 122″ for mapping biological signals, and a tip electrode 146for ablating or creating scar tissue to eliminate unwanted arrhythmiassuch as atrial fibrillation.

FIG. 12 is very similar to FIG. 5 except that it has a single shield 126which shields an L-C trap filter 136 which is connected between lead 114and lead 114′. In the art, this is known as a differential mode L-C trapfilter.

FIG. 13 is similar to FIG. 12 except that the L-C trap filter 136 hasbeen replaced by a general diverter element. In this case, the diverterelement 136 can be an L-C trap filter, a number of multi-element lowpass filters or single element capacitive filters. These are showncombined with impeder elements 132 and 134. The impeder element wouldtypically include an inductive component.

FIG. 14 illustrates the embodiment of FIG. 13 wherein the impederelements 132 and 134 are inductors.

FIG. 15 is very similar to FIG. 13 wherein the impeder elements arebandstop filters 128 and 130. These bandstop filters would normally betuned to be resonant at an MRI RF pulsed frequency or range offrequencies.

FIG. 16 is a graph showing impedance versus frequency for the idealparallel bandstop filter circuit 130 of FIG. 6 or 15. As one can see,using ideal (zero resistance) circuit components, the impedance measuredbetween the lead 114 and the tip electrode 122 is zero until oneapproaches the resonant frequency f_(r). At the frequency of resonance,these ideal components (the parallel inductor L and capacitor C)resonate together to approach an infinite impedance. In general, thefrequency of resonance f_(r) is given by the equation shown in FIG. 16and is selected to be the center frequency of an MRI RF pulsedfrequency. For example, for a 1.5 Tesla hydrogen scanner, the RF pulsedfrequency as determined by the Lamour equation is 42.56 times themagnetic field strength in Tesla. This is approximately 63.84. MHz.Accordingly, the resonant frequency of the bandstop filter 130 would beselected to be centered approximately around 63.84 MHz.

It should be noted that not all 1.5 Tesla MRI scanners have exactly thesame static magnetic field strength. This results in variation of the RFpulsed frequency by over ½ MHz. It is desirable that the bandstopfilters 128 and 130 provide substantial attenuation or 3 dB bandwidthover this entire range. Similar variations occur for other commonlylabeled MRI scanners, such as 3 Tesla scanners.

FIG. 17 is a drawing of a generic unipolar AMD 100 and lead 114, with abandstop filter 130 added at or near the distal electrode) 22. Theinductor L has a resistance element R_(L) in series with it. Thecapacitor C also has a resistance R_(C) in series with it. Theresistances R, and R_(C) can be separate discrete resistors or they arelosses of the inductor and capacitor elements themselves. In general,the resistance R, will be the resistance of the circuit traces or wiresused to form the inductor L. The capacitor C has ohmic losses R_(C) dueto the resistance of its internal electrode plates, connection to itselectrode plates, and dielectric losses. In the capacitor industry thisis known as the capacitor's equivalent series resistance or ESR. Thebandstop filter circuit 130 illustrated in FIG. 17 is a “real” bandstopfilter in that the resistive losses are included. This makes it distinctfrom the ideal bandstop filter circuit shown in FIGS. 6 and 15. Thepresence of the bandstop filter 130 will present a very high impedanceover a specific range of MRI RF pulsed frequencies to prevent currentsfrom circulating through the distal electrode 122 into body tissue atthis specific frequency range.

FIG. 18 is a family of curves 152, 154, 156 and 156′ which show theattenuation in dB versus frequency for the bandstop filters 128 and 130.Curve 152 represents the use of very high Q inductor and capacitorcomponents. If the capacitor and the inductor were ideal, meaning thatthey would both have zero resistance, then there would be no 3 dbbandwidth at all between points “a” and “b”. However, since in the realworld both the inductor and the capacitor do have losses, a 3 dbbandwidth separation between points “a” and “b” is achieved. It is veryimportant that there be suitable bandwidth for two reasons: one, the MRImachine has gradient fields which change the RF frequency. This is howthe MRI machine selects a slice to image, for example, through the knee.It does this by modifying (grading) the 1.5 Tesla or main static fieldby using a gradient field. This causes the Lamour frequency to change.Accordingly, one can see that some bandwidth is required centered aroundthe specified pulsed resonant frequency of the MRI equipment so that allof these frequencies are properly attenuated in an implanted lead. Ifone were to deliberately use an inductor with a very high DC resistanceand a capacitor with very high ESR, this would result in very low Qcomponents and the resulting attenuation curve 154. The low Qattenuation curve 154 attenuates over a very broad range of frequencies;however, the amount of attenuation in dB has been sacrificed. Ingeneral, in practice, it is easier to purchase monolithic ceramiccapacitors with relatively high Q values. Accordingly, the Q of theinductor L can be controlled by increasing R_(L). Accordingly, it is afeature of the present invention that the resistance of the inductor becontrolled to also control the overall Q and resulting 3 dB bandwidth ofthe parallel resonant bandstop filter 128, 130.

Attenuation curves 156 or 156′ shown in FIG. 18 are generally preferred.One can do this by controlling the relative Q of the inductor and thecapacitor components of the bandstop filter 128, 130. In one embodiment(156′), the Q of the inductor would be relatively low and the Q of thecapacitor would be relatively high. This means that the inductor wouldhave a relatively high internal resistance and the capacitor would havea relatively low equivalent series resistance. This is achieved by usingmultiple turns of relatively small wire to create a high DC resistancein the inductor, and by using multiple and robust electrode plates tokeep the equivalent series resistance (ESR) of the capacitor relativelylow. The overall Q of the bandstop filter is thus selected to balanceimpedance at the selected frequency versus frequency bandwidthcharacteristics. The values of the inductor and the capacitor selectedare such that the bandstop filter 128, 130 is resonant at a selectedfrequency, and preferably selected to attenuate current flow through thelead or electrode along a range of selected frequencies. Such afrequency or range of frequency may include an MRI RF pulsed frequency.Typically, the Q of the inductor is relatively low or moderate, and theQ of the capacitor is relatively high or moderate to select the overallQ of the bandstop filter. That is, the inductor has a relatively highresistive loss R_(L), and the capacitor has a relatively low equivalentseries resistance R_(C).

For medical implant applications it is very important that the implantedleads and their associated electrodes at the distal tips be very small.It is particularly important that the cross-sections or diameters of thebandstop filters be very small for easy endocardial insertion into thevenous system of the human body. The present invention meets thesecriteria by using a novel combination of components that aremechanically mounted in series, but whose lumped elements areelectrically in parallel. The components generally consist of commercialoff-the-shelf miniature chip capacitor and inductor components. Theseare generally manufactured in high volume throughout the world.Accordingly, they are very inexpensive, but more importantly, they arevery small in size. By way of example, twenty years ago a small sizedmonolithic chip capacitor (MLCC) would be 0603, meaning that it would be0.060 inch long by 0.030 inch in width. In comparison, current inductorand capacitor chip components can be purchased as small as 0201 or01005. This means that they are so small that they literally can fitthrough a pepper shaker. Human hands generally cannot assemblecomponents this small. Accordingly, micro-robotic manufacturing is thepreferred method of manufacturing the novel components assemblies of thepresent invention, wherein the components typically are delivered ontape and reel and fed into the robots which pick and place thecomponents and then go through a series of steps including additionalcomponent placement, wave soldering, cleaning, automatic opticalinspection and automated electrical testing. All of this is done in alinear robotic manufacturing operation that is completely or nearly freeof human hands. In cardiac rhythm applications (pacemakers and ICDs), adesirable lead size is 6 French (0.079 inches in diameter). For deepbrain stimulator applications, an even smaller size is desirable, suchas 3 French, which is 1 millimeter in diameter or 0.039 inches. US2007/0112398 A1 discloses a number of methods of manufacturing novelbandstop filters for placement in the lead systems of active implantablemedical devices. The present invention extends these concepts further.

In mammalian implant applications, the shielded inductors, diverters,impeders, and bandstop filters of the present invention should be smalland placed in series with the implanted lead or electrode of the medicaldevice. In general, the diameter is much more important than the volumeor length of the passive network package to be placed in series with animplanted lead 114, 114′ This is because leads are typically introducedinto the human body either by tunneling or transvenous insertion. Insuch applications, it is necessary that the shielded lead componentassembly be EMI shielded, biocompatible and highly reliable. Althoughcommercial off-the-shelf capacitor and inductor components are verysmall in size, arranging them such that they are electrically coupled inparallel can increase the size of the bandstop filter wherecomplications can arise in the placement and use of the implanted leador electrode.

Commercial off-the-shelf capacitor and inductor components are typicallynot entirely comprised of biocompatible materials. However, inaccordance with the present invention, the shielded inductor L andcapacitor C elements can be constructed to be completely biocompatible.In this case it would be not necessary to place them in a biocompatiblehermetic container. Just an open ended EMI shield would suffice. Thiswould have great advantages in further reducing both size and cost. Inthis regard, US 2009/0116167, U.S. Pat. No. 7,535,693, and US2009/0259265, are incorporated by reference.

With reference to FIG. 19, it is a feature of the present invention thatcustom or “off-the-shelf” non-biocompatible miniature inductor L andcapacitor C components are mechanically installed in shielded(conductive) hermetic packages or containers 158 in series, but haveelectrical circuit traces that couple the lumped inductor and capacitorelements electronically in parallel, thereby forming bandstop filters128, 130 as described above. FIG. 19 illustrates a hermetically sealedshielded container 158 having the inductor (L) and capacitor (C)components installed therein in series with one another, but whoselumped L and C elements are coupled electronically in parallel, so as toform one or more bandstop filters 128, 130. The shielded housing 160 forthe hermetically sealed container 158 is very small in diameter orcross-section and can be disposed between portions of an implantablelead 114, within an electrode assembly, etc.

FIG. 20 is a cross-sectional view taken generally along line 20-20 ofFIG. 19 and shows the various component parts of the shieldedhermetically sealed container 158. The shielded housing 160 can becomprised of a biocompatible metal or alloy, such as titanium, platinum,platinum-iridium, gold, palladium, tantalum, carbon, niobium, etc., oralloys thereof. The shielded housing 160 can also be a non-metallicmaterial, such as sapphire, ruby, lumina, ceramic, glass, etc, having athin layer of conductive metal deposited on either its inside or outsidesurface. For example, if the non-metallic shield was cylindrical, ametal coating could be applied to its outside diameter. In this case,the metal coating should be biocompatible and could be applied byelectroplating, sputtering, chemical vapor deposition, cladding, or thelike. The inductor L and the capacitor C are disposed on a substrate 162and physically arranged in series, or end-to-end with one another, yetconductively or electronically coupled to one another in parallel.Circuit traces 164 and 166 are conductively coupled to the inductor Land capacitor C of the bandstop filter 128, 130 and extend to conductiveterminals 168 and 170 of hermetic seal assemblies 172 and 174. Theconductive terminals 168 and 170 are designed to be conductively coupledto portions of the implantable lead 114, 114′ or electrode assembly.

FIG. 21 is an enlarged perspective view of the hermetic seal assembly174 from FIG. 20, having the terminal 170 extending therethrough to acrimp, solder joint or laser weld tip 176. The electrical connection tothe tip 176 could also be formed by thermal-setting conductiveadhesives. The terminal 170 is attached to an insulator 178, which is inturn attached to an outer ferrule 180.

FIG. 22 is a cross-section drawing taken along line 22-22 from FIG. 20.The terminal 170 is preferably of a common platinum-iridium alloy, suchas 9010 or 8020. However, any biocompatible and suitable material couldbe used in place of platinum-iridium. Gold braze 182 forms a hermeticseal between terminal 170 and insulator 178. The insulator 178 may be apolished sapphire, ruby, polycrystalline alumina, or even glass or ageneral ceramic material. Sputtering would first be deposited on thesurfaces so that the gold braze 182 will readily adhere and wet. Goldbraze 184 forms a hermetic seal between insulator 178 and the ferrule180. Gold brazes 182 and 184 are generally pure gold brazes forbiocompatibility and long term reliability. The surface preparationprocess for the ceramic insulator 178 can be as follows: C-Axis singlecrystal, polycrystalline alumina (Al2O3), Zirconia Stabilized Aluminaand/or Yttria Tetragonal Zirconia Polycrystalline YTZP is etched usingRF plasma before PVD sputtering using a biologically compatible metallicsystem. Plasma cleaning removes organic surface contamination andhydroxyl/oxides resulting in a higher energy surface. This activatedsurface readily forms strong covalent bonds with metallization atomspromoting robust, hermetic adhesion. Through industry standard processrefinements, the resulting low stress, dense coating does not spall offor blister and improves the function and reliability of the final brazedjoint. The outer ferrule 180 is also, preferably, of platinum-iridiumsince it's very easy to laser weld. It is also radio-opaque.

In the preferred embodiment, the insulator 178 would be a polishedsapphire. It would then go through a plasma-etch process, such as a 500watt plasma-etch, to increase its surface roughness. Titanium-molybdenumor niobium metallization would be a preferred sputter material foradhesion and wetting of the associated gold braze pre-forms.

In FIGS. 20 and 21, one can see that the interior tip 176 of theterminals 168 and 170 has been extruded to be fitted into an aperture,socket, etc. of the conductive substrate or circuit traces 164 and 166.Alternatively, the interior tip 176 may have an aperture therethrough sothat a crimped connection can be formed between it and the conductivesubstrate or circuit traces 164 and 166, and subsequently laser welded.The method of attachment to the interior tip 176 will vary in accordancewith the type of attachment desired to the internal circuitry of thebandstop filter 128, 130. In any event, the conductive terminals 168 and170 are conductively coupled to the bandstop filter 128, 130 as theassociated hermetic seal assemblies 172 and 174 are slid into place andhermetically sealed by laser welding 186 to the housing 160 of thecontainer 158. FIG. 23 is an end view taken along line 23-23 in FIG. 20.

Again, FIG. 20 shows the bandstop filter 128 and 130 comprised of theinductor L and capacitor C, and the flexible circuit substrates 164 and166 extending therefrom, attached to the terminals 168 and 170 so as toplace the terminals 168 and 170 in electrical series with one another.However, the inductor L and the capacitor C, although placed end-to-endand physically in series with one another, are conductively coupledelectrically with one another in parallel. An insulating material 188,such as a thermal-setting non-conductive polymer, at least partiallyfills the remainder of the EMI shield housing 160 to provide protectionand mechanical robustness to the overall container assembly 158. Thisstructure lends itself to a novel “ship-in-the-bottle” method ofmanufacturing. That is, all of the elements contained within the shieldhousing 160 are pre-assembled outside the housing. In particular, theterminal 168, the substrate 162 containing the inductor L and capacitorC, and the opposite terminal 170 and the associated hermetic seals 172and 174, are all pre-assembled outside of the overall EMI shield housing160. This facilitates proper electrical connections and electricaltesting of the pre-assembly. In addition, this entire subassembly can gothrough high reliability screening. Typically, this would consist ofthermal cycling or thermal shock followed by a burn-in, which meansapplying a relatively high voltage at elevated temperature to thecircuit components and then comprehensive electrical test afterwards.Once all of this has been done, this entire pre-assembly is slippedinside the overall cylindrical EMI shield housing 160 and then a finallaser weld 186 is formed.

FIG. 20 also shows an optional conformal coating 190 which is providedover the two gold brazes 182 and 184. This conformal coating 190 couldalso be applied to the entire outer surface of the housing 160 and aportion of terminals 168 and 170, as well as optionally over theelectrical attachments to the lead system. This conformal coating 190 isimportant to provide electrical isolation between the two terminals 168and 170. When directly exposed to body fluids (which containelectrolytes), gold can migrate in the presence of a voltage bias. Ithas been shown that pacemaker pacing pulses in saline solution canactually cause a gold electro-migration or electroplating action. Theconcern is that the gold braze materials 182 and/or 184, under voltageor pulse bias, may over time migrate or deposit (electro-plate) ontoanother surface such as the terminal 170 or the housing 160, which couldnegatively affect the long-term hermeticity and reliability of thehermetic seal assembly 174. Accordingly, the conformal coating orbackfill 190 is placed as shown to cover both of the gold brazes 182 and184. The conformal coating 190 may comprise thermal-settingnon-conductive adhesives, silicones, parylene (which is vapordeposited), and the like, including epoxies, polyimides, polyethyleneoxide, polyurethane, silicone, polyesters, polycarbonate, polyethylene,polyvinyl chloride, polypropylene, methylacrylate, para-xylylene, andpolypyrrhol. In particular, Epo-tek H81 is considered a preferred epoxywhich has already been tested for long-term biocompatibility. Theimportance of providing electrical isolation across components, such asbandstop filters, is more thoroughly described in U.S. patentapplication Ser. No. 12/873,862 which is incorporated herein byreference.

A complete conformal coating 190 over the entire shield housing 160 maybe desirable to provide electrical isolation between the conductiveterminal pins 168 and 170. This provides critical performance capabilityin the event of complete saturation of the housing 160 in saline orbiological fluid. Additional performance benefits for a conformalcoating 190 include lubricity, radiopacity, and wear resistance.

FIG. 24 illustrates a multi-layer substrate or flex cable 192 onto whichthe inductor L and capacitor C are mounted. The inductor L is a chipinductor having first and second conductive termination surfaces 194 and196 which are spaced from one another in non-conductive relation. Thecapacitor C also has first and second conductive termination surfaces198 and 200 which are spaced apart from one another in non-conductiverelation. The chip inductor L can be any number of chip inductor types,however the present invention is also not limited to chip inductorsonly. The inductor L could also be a solenoid inductor, a toroidalinductor, or any type of inductor that is known in the prior art.Moreover, the chip capacitor C can be any number of chip capacitortypes, but the present invention is not limited to chip capacitors only.The capacitor C may be of many different types of capacitortechnologies, including film capacitors, tantalum capacitors, monolithicceramic capacitors, electrolytic capacitors, feedthrough-typecapacitors, or even tubular capacitors. FIGS. 24 and 25 show that theinductor L and the capacitor C are physically disposed in seriesrelative to one another, such that they are generally aligned with oneanother along a common longitudinal axis and placed end-to-end. However,as shown in FIGS. 25 and 26, the inductor L and the capacitor C areconductively or electrically coupled to one another in parallel. FIG. 26is an electrical schematic diagram of the bandstop filter 128, 130 ofFIGS. 24 and 25. For a more complete description of how to disposeimplantable lead components physically in series but electrically inparallel, reference is made to US 2010/0100164 A1, the contents of whichare incorporated herein.

FIGS. 27-29 illustrate a configuration where a chip inductor L isphysically disposed in series with a feedthrough capacitor C, and yet iselectrically connected in parallel to form a bandstop filter 130. Thechip inductor L and the feedthrough capacitor C are disposed within anEMI shielded hermetic container 158 comprising a conductive housing 160of a biocompatible material which includes one open end, and a hermeticseal assembly 174 disposed within the open end of the housing 160. Theconductive terminal 168 is conductively coupled to the housing 160 by alaser weld 202. The first conductive termination surface 194 of theinductor L is conductively coupled to the housing 160 by means of asolder, braze, or conductive adhesive 204 or the like. The secondconductive termination surface 196 of the inductor L is similarlyconductively coupled by means of a solder, braze, or conductive adhesive206 or the like, to a conductive bracket 208 which is also conductivelycoupled to an extension 210 of the conductive terminal 170 which extendsthrough a central passageway of the feedthrough capacitor C. The firstconductive termination surface 198 of the capacitor C is conductivelycoupled to the housing 160 by means of conductive adhesive 212 or thelike, and the second conductive termination surface 200 of thefeedthrough capacitor C is conductively coupled to the extension 210 ofthe conductive terminal 170 by means of conductive adhesive 214 or thelike. The hermetic seal assembly 174 disposed within the opening to thehousing 160, and which prevents direct contact between body fluids andthe inductor L, the capacitor C and related electrical components, isessentially the same as the hermetic seal assembly 174 illustrated inFIGS. 20-23. The illustrated structure advantageously eliminates onehermetic seal assembly in comparison with previously illustratedembodiments, by providing a terminal 168 which is shorted to theconductive EMI shield housing 160. As shown, the optional conformalcoating 190 is applied over the entire outer surface of the housing 160as well as a portion of the terminals 168 and 170. This conformalcoating 190 advantageously provides additional electrical isolationbetween the two terminals 168 and 170.

FIG. 30 illustrates a solenoid inductor 216 wrapped around anon-ferromagnetic core 218. The term solenoid inductor as defined hereinincludes any inductor geometry whose magnetic fields 219 are alignedgenerally along the central axis of the lead and/or the shield 126. Theinductor 216 consists of coils of wire 220 which can be single ormulti-layer as illustrated in cross-section in FIG. 32. The inductor 216may be wound on a magnetic material such as a ferrite core, however,this is highly undesirable for MRI applications. This is because the MRImain static field would tend to saturate such high permittivity (k)ferrite materials. Accordingly, the shielded inductors of the presentinvention are generally comprised of air or non-ferromagnetic materialsas shown in FIG. 31, where the inductor wires 220 are coiled around asupport structure 222 of a non-magnetic material, such as a ceramic orplastic. This makes the coil 224 of FIG. 31 equivalent to an “°air-wound” inductor. These so-called air coils are not veryvolumetrically efficient and tend to have a magnetic field 219 asillustrated in FIG. 32. This is well known in physics and for a DC case,would generate a north and south pole. In an AC case, which is the casefor an MRI RF application, these field lines would be alternating at theRF frequency of the MRI RF pulsed field. These field lines 219 wouldtherefore build-up and collapse which also reverses the induced currentsagain at the frequency of the RF pulsed field. The inventors havedetermined that the field lines 219 of the solenoid inductors 216, 224are affected when the inductor is placed inside of an electromagneticinterference shield 126. As previously described in connection withFIGS. 20 and 27, this electromagnetic interference shield 126 can alsobe the housing 160 of a hermetically sealed container.

As one can see in FIG. 33, when the solenoid inductor 216, 224 is placedwithin a conductive shield 126, the magnetic field lines 219 of thesolenoid inductor coil 220 tend to capture and induce currents in theshield 126 which affects the energy stored in the coil's magnetic field,and therefore the inductance value of the inductor 216, 224. FIG. 33illustrates a worst case for a solenoid coil's magnetic fields 219wherein the shield 126 has a very high permeability. The highpermeability of the shield 126 creates a low reluctance path for themagnetic fields 219 which tends to capture some amount of the coil'smagnetic field 219 in the shield 126. The amount of flux captured by theshield 126 is directly related to the material's permeability; as thepermeability increases, more magnetic field lines 219 tend to becaptured within the shield. In a preferred embodiment, the shield 126 isof a biocompatible material such as platinum-iridium alloy which has arelatively low permeability. Accordingly, for a shield 126 ofplatinum-Iridium (or equivalent biocompatible metals such as titanium,stainless steel, niobium), the magnetic field lines 219 do notcompletely collapse into the shield walls as shown in FIG. 33, butrather the field lines 219 penetrate and propagate outside the shield126. For both high permeability and lower permeability shields 126, thevalue of the solenoid coil inductance in nanoHenries or microHenries isshifted when one measures this value with the inductor coil 220 outsideof the shield 126 (in air) as opposed to inserting the inductor coil 220into the shield. When the inductor 216, 224 is a component of an LCbandstop filter 128, 130, it is very critical that this change ininductance be accounted for in the design. If it is not properlyaccounted for, the resulting resonant frequency of the L-C bandstopfilter 128, 1.30 may not be centered on an MRI band of RF pulsedfrequencies.

FIG. 34 illustrates a prior art thick film or chip inductor 226 takenfrom U.S. Pat. No. 5,970,604, the contents of which are incorporated byreference. As defined herein, the term “chip inductor” includes anyinductor winding or circuit trace geometry whose magnetic fields aregenerally aligned at 90 degrees to the central axis of the lead 114and/or the shield 126. Such chip or thick film inductors 226 may beutilized in connection with the present invention, either alone or inconnection with a parallel capacitor C to form a bandstop filter 128,130.

FIG. 35 shows an exemplary arrangement of circuit traces 228 that formthe thick film inductor 226. The electromagnetic field lines of such aninductor are generally opposite orthogonally to those for the solenoidinductors of FIGS. 30 and 32.

FIG. 36 illustrates the ideal (in air) electromagnetic fields 230 aroundthe thick film inductor 226 of FIGS. 34 and 35. In this case, themagnetic fields 230 are directed at 90 degrees to the direction of theimplanted lead center line. When this type of thick film inductor 226 isinserted into a shielded housing 126, the field lines will inducecurrents into the surrounding electromagnetic shield as shown in FIGS.37 and 38.

FIGS. 37 and 38 illustrate how the magnetic fields 230 of the chipinductor of FIGS. 34 and 35 tend to be captured and induce currents intothe shield walls 126 which affects the energy stored in the coil'smagnetic field and the coil inductance value. FIG. 37 is a worst casefor a chip inductor's magnetic fields 230 wherein the shield 126 has avery high permeability. The high permeability of the shield creates alow reluctance path for the magnetic fields 230 of the inductor 226which tends to capture a great deal of the coil's magnetic field 230 inthe shield walls 126. FIG. 38 comes from electromagnetic field modelingand shows the magnetic field pattern 230 of the chip inductor 226 whichis orthogonal to the both central axis of the shield 126 and the leadcentral axis. In a preferred embodiment, the shield 126 is of abiocompatible material such as platinum-iridium alloy which has arelatively low permeability. Accordingly, for a shield 126 ofplatinum-Iridium material (or equivalent biocompatible metals such astitanium, stainless steel, niobium), the magnetic field lines 230 do notcompletely collapse into the shield walls 126 as shown in FIG. 37, butrather the field lines 230 penetrate and propagate outside the shield126. For a chip inductor 226 inserted inside of either a highpermeability or lower permeability shield 126, the value of theinductance in nanoHenries or microHenries is different if one measuresthis value with the inductor outside of the shield 126 (in air) asopposed to inserting the inductor into the shield. This shift is lessthan that for a solenoid inductor 216, 224, but is still significant.Still, when the chip inductor 226 is a component of an L-C bandstopfilter 128, 130, it is very critical that this change in inductance beaccounted for in the design. If it is not properly accounted for, theresulting resonant frequency of the L-C bandstop filter 128, 130 may notbe centered on an MRI band of RF pulsed frequencies which would make itineffective.

In general, the amount of induced current from the magnetic field 230 ofa chip inductor 226 in the shield 126 encompasses less area and lessmagnitude as compared to the solenoid inductors 216, 224 of FIGS. 31-33.In other words, there is less energy loss and inductive shift from achip inductor geometry as compared to a solenoid inductor type ofarrangement.

FIG. 39 is a graph of histograms from various prototypes of solenoidinductors that were built as a component of L-C resonant bandstopfilters. A network analyzer was used to measure the resonant frequencyof the bandstop filter within the limit of the component tolerances ofthe inductors and capacitors. The left hand histogram is a graph of theL-C bandstop filter resonant frequencies measured with the bandstopfilter disposed well outside of the overall EMI shield 126 (in otherwords, in air). The mean resonant frequency was measured to be 59.29 MHzwith a standard deviation of 0.2486 based on 71 units measured. Then thebandstop filters were placed inside of a shielded housing 126 and againtheir resonant frequencies were measured. In this case the mean resonantfrequency was determined to be 63.65 MHz with a standard deviation of0.1958 with a total of 100 samples measured. In all cases, thecapacitors were 15.9 picofarad and mounted on circuit boards similar tothose shown in FIG. 24. In this case, the shielded housing was made ofplatinum-iridium. Remarkably, this effect on the inductor fieldsaccounts for a shift in resonant frequency of the bandstop filter of 4.3MHz or 6.8%. The frequency of resonance f_(r) for a bandstop filter isgiven in FIG. 16 where one can see there is an inverse relationshipbetween the square root of the inductance and capacitance and theresonant frequency. Assuming the capacitance is held constant at 15.9picofarads and solving the resonant frequency equation for inductance,this means that the inductance in air on average was 453.2 nanoHenriesand dropped to an effective 394.3 nanoHenries when inserted into thesurrounding EMI shield housing 126. This is an average shift of 58.9 orapproximately 59 nanoHenries, which is about a 13% shift in theinductance value. Accordingly, in order for the L-C bandstop filter tobe properly resonant in an MRI RF pulsed center frequency, this shift inthe inductance in air versus insertion into the MRI shield 126 must beproperly accounted for. There are many variables that come into play inthis calculation, including the physical properties of the inductor, itsorientation as a solenoid or a chip inductor, the thickness and diameterof the surrounding electromagnetic shield and its high frequencymaterial properties, including its high frequency resistance.

FIG. 40 illustrates impedance versus frequency curves for circuit boardswith either thick film chip or wire wound or solenoid inductors (thecentral axis and corresponding magnetic fields of a typical thick filminductor are oriented at 90 degrees to a typical wire wound solenoidinductor). This graph demonstrates the effect of placing the thick filmor wire wound inductor inside a cylindrical metal shield tube 126.Notice that the solenoid-type wire wound inductor boards exhibit asignificant shift in resonant frequency and impedance. Having a veryhigh impedance at resonance is desirable to prevent undesirable MRI RFinduced currents from flowing into surrounding body tissues.

FIG. 41 is a PSPICE computer model developed by the inventors to be ableto predict the resonant circuit behavior of a solenoid inductor in air.This model was modified as shown in FIG. 42 to add mutual inductancecoupling to a surrounding electromagnetic shield 126. This PSPICE modelis used in conjunction with the present invention to predict that amountof resonant shift and the amount of inductive offset one needs to makewhile designing inductors for shielded L-C bandstop filters.

FIG. 43 is the frequency response predicted by the PSPICE models for theinductor inside and outside of a platinum-iridium shielded housing. Asone can see, the PSPICE model very accurately fits the empirical datapreviously plotted in FIG. 40.

The PSPICE model can be used to adjust a first inductive value with theinductor outside of a shielded housing so that a second inductive valuewith the inductor inside of a shielded housing has the proper value. Forexample, in a resonant tuned L-C bandstop filter it is very importantthat the inductor value and the capacitor value have a fairly tighttolerance so that the resulting resonant frequency occurs in the centerof a range of MRI RF pulsed frequencies.

With reference now to FIG. 44, a passive electrode fixation tip 232typically used in cardiac pacemaker applications is shown in which theshielded inductor, passive network or bandstop filter assembly of thepresent invention can be incorporated.

FIG. 45 is a sectional view of a portion of the passive electrode 232taken along the line 45-45 from FIG. 44, and illustrates a hermeticallysealed package consisting of a passive distal tip electrode 122 which isdesigned to be in intimate contact with body tissue, such as inside theright atrium of the heart. A hermetic seal is formed at laser weld 234as shown between the tip electrode 122 and a metallic ring 236. Goldbrazes 238 are used to separate the metallic ring 236 from the shieldsurface 126 by use of an intervening insulator 240. This insulator 240could typically be of alumina ceramic, other types of ceramic, glass,sapphire or the like. The 126, which also acts as an energy dissipatingsurface EDS, is typically gold brazed to the other side of the insulator240 as shown. An inductor L, such as an inductor chip is shown connectedbetween the distal tip electrode 122 and a terminal pin 170 which isattached as by laser welds 244 to the end of the lead 114 extendingthrough the body to the AMD. As shown, terminal pin 170 protrudesthrough a hermetic seal assembly 174.

The shield/energy dissipating surface 126 of FIG. 45 is typically of abiocompatible metal, such as titanium, platinum or the like. It isimportant that the shield/energy dissipating surface 126 be bothelectrically conductive and thermally conductive so that it can transferRF and thermal energy into body fluid or tissue. The shield/energydissipating surface 126 can be roughened or even corrugated or bellowedto increase its surface area and therefore its energy dissipatingproperties into surrounding body fluids or body tissue.

Capacitive elements C and C′ shown in FIG. 45 are designed to act as alow impedance at higher frequencies. Electrical connections 246 couplethe capacitor C to the shield/energy dissipating surface 126. This formsa broadband low pass filter wherein the inductor L acts in cooperationwith the capacitive elements C and The presence of the inductor Lenhances the performance of the capacitor elements C and C′, which aretypical off-the-shelf commercial monolithic ceramic capacitors (MLCCs)such as those illustrated in FIGS. 47 and 48.

An advantage in using a capacitor C as a selective frequency element isthat it tends to act as a broadband filter which will attenuate a rangeof MRI frequencies. For example, placement of an effective capacitor Ccould attenuate 64 megahertz, 128 megahertz and higher MRI frequencies.However, if one were to use an L-C series trap filter as shown in FIG.8, then this would only be effective at one MRI frequency, for example64 megahertz only. Of course, as already been disclosed herein, onecould use multiple L-C trap filters. However, in a preferred embodimentthe use of a capacitor as is desirable because with a two-element L-typelow pass filter, one can attenuate a broad range of MRI RF pulsedfrequencies.

The schematic diagram for the circuitry of FIG. 45 is shown in FIG. 46.Capacitors C and C″ are actually in parallel and act as a singlecapacitive element. The reason for multiple capacitors is to obtain ahigh enough total capacitance value so that the capacitive reactance isvery low at the frequency of interest (for example, 64 MHz for a 1.5 TMR system).

An alternative capacitor C″ for use in the circuit of FIG. 45, known asa unipolar feedthrough capacitor, is shown in FIG. 48. It has outsidediameter and inside diameter termination surfaces and for electricalcontact. Feedthrough capacitors can be unipolar or multipolar. These arecompletely described in the prior art; for example, refer to U.S. Pat.No. 7,363,090, U.S. Pat. No. 4,424,551; U.S. Pat. No. 5,333,095; andU.S. Pat. No. 6,765,779.

FIG. 49 is similar to FIG. 45 except that the inductor element L is wirewound around a non-ferromagnetic mandrel 222 (formed from a materialsuch as a ceramic or plastic). This type of solenoid wound inductor Lhas much higher current handling capability as compared to the inductorchip of FIG. 45. The inductor chip of FIG. 45 can be fabricated from avariety of shapes including Wheeler spirals, thick film inductors, andthe like. It is important that the inductor element L be able to handlesubstantially high currents when it is in series with the lead 114. Thereason for this has to do with either ICD applications for shockelectrodes or automatic external defibrillation (AED) events. AEDs havebecome very popular in government buildings, hospitals, hotels, and manyother public places. When the external defibrillator paddles are placedover the chest of a cardiac pacemaker patient, the high voltage thatpropagates through body tissue can induce powerful currents in implantedleads. Accordingly, the inductor L has to be designed to handle fairlyhigh current (as high as the 4 to 8 amp range in short bursts). The wirewound inductor L of FIG. 49 has wire of a larger cross-sectional areaand is therefore a higher current handling inductor.

FIG. 50 illustrates an entirely different approach for the diverting ofRF energy away from the electrode tip 122 to the shield/energydissipation surface 126. Shown is an electrical connection 248 between afirst inductor L and the distal tip electrode assembly 122. The otherend of the first inductor L is connected to a second inductor L′ whichis in turn electrically connected at 250 to the hermetic terminal pin170. The capacitor C is connected between the junction of the twoinductors L and L′ at electrical connection 252. The other end of thecapacitor is electrically connected to the shield energy dissipatingsurface 126. An insulating sleeve (not shown) can be used to ensure thatthe capacitor termination and electrical connection 252 does notinadvertently make contact (short out) with the shield/energydissipating surface 126.

The electrical schematic for FIG. 50 is shown in FIG. 51. This forms alow pass filter (in this example, a T filter), which tends to enhancethe filtering performance by directing more of the RF energy to theshield/energy dissipating surface 126. As previously mentioned, a singleor multi-element low pass filter would attenuate a broad range of MRIfrequencies and would be an advantage in the present invention for thatreason. In accordance with the present invention, it is important thatthe value of the inductance for either the chip inductor L of FIG. 45,the solenoid inductor L of FIG. 49, or the chip inductors L, L′ of FIG.50 have their first inductive values adjusted so that their inductance,when inserted into the overall shield/energy dissipating surface 126, sothat the resultant package value is correct.

FIGS. 52 and 53 show a generic prior art active fixation distal tipelectrode 254 which is typically used in conjunction with cardiacpacemakers. There is a metallic housing 256 which contains a sharptipped distal helix coil 258. In FIG. 53, this helix coil 258 is shownin its retracted position, which enables the physician to insert thefixation tip assembly 254 endocardially through the venous system,through the atrium, and through the tricuspid valve into the rightventricle so it does not snag or tear any tissue, and is designed to beextended and screwed into myocardial tissue. Once it is in theappropriate position, the physician then turns leadwire spline assembly260 in a clockwise rotation. This is done outside the pectoral pocketwith the lead 114 protruding from the body. A torque tool is generallyapplied so that the physician can twist or screw the helix coil 258 intoplace. Protrusion 262 acts as a gear so that as helix coil 258 isturned, it is screwed forward. This makes for a very reliable fixationinto myocardial tissue. The helix coil 258 is generally attached by alaser weld 264 to an end of the spline assembly 260 as shown. Attachedto spline assembly 260, usually by laser welding, is the lead 114 comingfrom the AMD. An optional feature 266 is placed on spline assembly 260to create a positive stop as the physician is turning the leadwireassembly and screwing the helix coil 258 into body tissue. Of course,all of the materials of the active fixation tip 254 shown in FIG. 53 arebiocompatible. Typically, the helix coil 258 is made of platinum iridiumalloy and would be coated with various materials to improve electricalperformance. The housing 256 would generally be composed of titanium oranother equivalent biocompatible alloy. The spline 260 is generally aplatinum iridium alloy.

FIG. 54 illustrates applying a shielded inductor L or a bandstop filter130 to the active fixation distal tip 254 shown in FIG. 53. One can seethe attachment from the metallization 268 of inductor L or bandstopfilter 130 shown attached to the spline 260. This is typicallyaccomplished by a gold braze preform 270. In this case, the spline 260has been counter-bored to receive the end of inductor L or bandstopfilter 130. This allows the gold braze material 270 to angle up alongthe sides of the assembly, thereby adding shear strength. A similar goldbraze preform 272 is used to attach a distal tip helix pedestal 274 tothe metallization 276 of the inductor or passive network assembly. Ofparticular advantage is that the assembly illustrated in FIG. 54 can beconstructed entirely of low k, very high strength ceramics. In thiscase, pure alumina or porcelain would be preferred embodiments. Thesehave the advantage of being mechanically very rugged and also veryrugged to thermal shock such that it would take pure gold brazing. Byuse of all biocompatible materials, the assembly is greatly simplifiedin that it need not be hermetic. It would also be possible to replacethe gold brazes 270 and 272 with equivalent laser welds. One can seethat the end cap 278 has been modified in a novel way such to make itflush with the outside diameter of the housing 256. This allows one toincrease the inside diameter allowing room for the counterbore in thespline assembly 260. The metallic end cap 278 has been stepped so thatit is seated for convenient fixturing. The overall housing 256 for thetranslatable helix assembly. 254 is conductive and forms a shield inaccordance with the present invention around either the inductor L orthe bandstop filter 130. Of course, the inductor L or bandstop filter130 could be replaced with any low or high pass filter and/or activeelectronic circuit.

FIG. 55 is an adaptation of the generic prior art active fixation distaltip 254 illustrated in FIG. 54. The design allows: 1) body fluid tofreely penetrate to all surfaces interior to the active fixation distaltip 254; and 2) torque experienced by the helix 258 is not transmittedto any electronic component, such as the hermetically sealed bandstopfilter 130 and its associated electrical and mechanical connections. Thespline shaft 260 has been modified such that it has a relatively long,hollow cylindrical cup portion 280 which allows for installation of theinductor L or bandstop filter 130 inside of it. As will be seen, thiswill offer a number of important mechanical and biocompatibilityadvantages. The inductor coil 282 has either been wound around a mandrelwhich has been removed or is wound around a mandrel which isnon-ferromagnetic. In the preferred embodiment, the coil 282 is freestanding and is then backfilled with an insulative dielectric material284. The dielectric insulating material 284 is preferably dispensed as athermal-setting liquid. After curing at high temperature, the insulatingmaterial 284 is cured to form a solid. This material can be athermal-setting non-conductive epoxy or polyimide or the like. Analternative (not shown) would be to insert a rigid insulating sleevearound the inductor, which has a pre-formed shape. This could be used incombination with insulated inductive wire turns to control the seriesand parallel parasitic capacitance. The space in between the turns ofthe coil 282 and its relationship to the cup assembly 280 is importantas parasitic capacitances Cp and Cs are developed.

This arrangement is best understood by looking at the equivalentcross-sectional schematic diagram illustrated in FIG. 56. One can seethat there are parasitic capacitances Cs formed between the coil turnsand also parasitic capacitances Cp to the outer shield housing cupassembly 280. As shown in the schematic in FIG. 57, all of thesecapacitances add up to form a capacitance in parallel with the inductorL. Once the schematic of FIG. 57 is simplified, it becomes a shieldedparallel resonant bandstop filter 130 as shown in FIG. 58.

An additional advantage of having the inductor L or capacitor-inductor130 inside the housing 256 of the active fixation tip 254 is that thisprovides a substantial degree of protection to these delicate electroniccomponents. Doctors and other medical personnel are often notorious inthe way they handle lead systems. Things can get dropped, moved orplaced against them.

FIG. 59 illustrates a reinforced polyimide tubing 286. The typicalconstruction consists of a substrate layer 288, a braided or coiledmetallic shield layer 290 and an exterior layer 292 (see cross-sectionFIG. 60). The substrate 288 and exterior layer 292 are insulativewherein the embedded braided or coiled layer 290 is a conductive metal.In a particularly preferred embodiment, the insulative exterior layer292 would be eliminated such that the conductive shield 290 would be indirect contact with body fluid. Since the conductive shield 290 has arelatively very large surface area, RF energy can be conducted in thebody tissues without resulting in significant temperature rise. This isfurther described in US 2010/0160997 A1 and US 2010/002300 A1, both ofwhich are herein incorporated by reference. The most common braid coil290 material is 304V stainless steel. Other metallic materials can alsobe used. The embedded braid coil 290 accomplishes RF shielding inaccordance with the present invention. FEP and PTFE coatings can beadded to the outside diameter both to enhance slickness (lubrication) tomake it easy to insert the lead into the body tissues.

FIG. 61 illustrates an alternative embodiment wherein an insulation tube294 is slipped over the lead 114. Then, a shield layer 290, such as aplatinum-iridium, is slipped over the insulation tube 294 as shown.

FIGS. 62 and 63 are similar to FIG. 61 except that the metal shield tube290 is replaced by wound wire strands 296 or wrapped foil 298,respectively, or other equivalent materials which are commonly used inshielded cables worldwide.

FIG. 64 shows an open mesh cross braided shield wire 290 instead of awound shield wire. The cross braid shield 290 is shown in more detail inFIG. 65, wherein one can see how the braided wires 300 interweave.

The thickness of the conductive shield 290 may require precise control.Thin deposition methods are capable of applying films in the nanometerrange. The skin depth or effective skin depth, due to limitedconductivity from surface scattering and such, of these thin films maybe of a thickness that external electromagnetic waves are not fullyattenuated.

Most applications will require full or near-full attenuation to preventsignificant currents on the internal sensitive components orconnections. However it may be desirable that the energy is not fullyattenuated, for example when it is desired to limit the amount ofcurrent needed to fully attenuate the incident electromagnetic wave toprevent over-heating. Further, multiple shields may be utilized toprevent overheating or allow limited energy to be attenuated on theinternal components to allow monitoring of the external environment forapplications such as automatic mode switching or data-logging.

Accordingly, from the foregoing it will be appreciated that the presentinvention resides in a shielded component or network for an activemedical device (AMD) implantable lead which has a length extending froma proximal end to a distal end, all external of an AMD housing. Apassive component or network is disposed somewhere along the length ofthe implantable lead, the passive component or network including atleast one inductive component having a first inductive value. Anelectromagnetic shield substantially surrounds the inductive componentor the passive network. Importantly, the first inductive value of theinductive component is adjusted to account for a shift in its inductanceto a second inductive value when shielded.

The inductive component may comprise a simple inductor, a low passfilter, an L-C trap, or a bandstop filter. When a bandstop filter or L-Ctrap filter is provided, the capacitive and inductive components aretuned to impede induced current flow through the implantable lead at aselected center frequency or range of frequencies, technically an MRI RFpulsed frequency or range of RF pulsed frequencies.

Although several embodiments have been described in detail for purposesof illustration, various modifications may be made without departingfrom the scope and spirit of the invention. Accordingly, the inventionis not to be limited, except as by the appended claims.

What is claimed is:
 1. An implantable lead, comprising: a) a lead bodyhaving an axial length extending from a proximal lead end to a distallead end, where the proximal lead end is configured to be permanently orremovably connectable to an active implantable medical device; b) a tipelectrode configured to be contactable with biological cells, where thetip electrode is disposed at or adjacent to the distal lead end; c) afirst conductor disposed within the lead body, the first conductorelectrically coupled to the tip electrode and extending to or adjacentto the proximal lead end, wherein the first conductor comprises aninsulated and self-resonant first inductive coil, where the firstinductive coil is disposed physically and electrically in series withthe first conductor, and where the first inductive coil is disposedproximal to the tip electrode in relation to the axial length of thelead body, wherein the first inductive coil comprises a first parasiticcapacitance between its adjacent turns creating a first totalcapacitance wherein a first equivalent electrical circuit schematiccomprises a first inductance in parallel with the first totalcapacitance and wherein the first inductive coil is self-resonant at ornear an MRI RF pulsed frequency and presents a high impedance in serieswith the first conductor providing at least 10 dB of attenuation at itsresonant center frequency, wherein the first inductive coil has a firstoverall circuit Q wherein the resultant 3 dB bandwidth is at least 100kHz; d) a ring electrode configured to be contactable with biologicalcells, where the ring electrode is disposed near the distal lead end andproximal to the first inductive coil in relation to the axial length ofthe lead body; e) a second conductor disposed within the lead body andinsulated from the first conductor, the second conductor electricallycoupled to the ring electrode and extending to or adjacent to theproximal lead end, wherein the second conductor comprises an insulatedand self-resonant second inductive coil, where the second inductive coilis disposed physically and electrically in series with the secondconductor, and where the second inductive coil is disposed proximal tothe ring electrode in relation to the axial length of the lead body,wherein the second inductive coil comprises a second parasiticcapacitance between its adjacent turns creating a second totalcapacitance wherein a second equivalent electrical circuit schematiccomprises a second inductance in parallel with the second totalcapacitance and wherein the second inductive coil is self-resonant at ornear an MRI RF pulsed frequency and presents a high impedance in serieswith the second conductor providing at least 10 dB of attenuation at itsresonant center frequency, wherein the second inductive coil has asecond overall circuit Q wherein the resultant 3 dB bandwidth is atleast 100 kHz; and f) a first electromagnetic shield at least partiallysurrounding the first inductive coil.
 2. The implantable lead of claim1, wherein the tip electrode is selected from the group consisting of apassive fixation tip electrode, a translational active fixation helixtip electrode, a ring electrode, a pad electrode, a deep brainelectrode, a spinal cord electrode, a catheter electrode and a cochlearelectrode.
 3. The implantable lead of claim 1, wherein the first totalcapacitance also comprises a parasitic capacitance between the firstinductive coil and the first electromagnetic shield.
 4. The implantablelead of claim 1, wherein the first electromagnetic shield is configuredto be electrically coupled to a housing of the active implantablemedical device.
 5. The implantable lead of claim 1, including a groundconductor electrically coupled at its one end to the firstelectromagnetic shield and wherein the other end of the ground conductoris configured to be electrically coupled to a housing of the activeimplantable medical device.
 6. The implantable lead of claim 1, whereinthe first electromagnetic shield comprises a plurality ofelectromagnetic shields.
 7. The implantable lead of claim 6, wherein theplurality of electromagnetic shields are electrically coupled to oneanother.
 8. The implantable lead of claim 7, including a groundconductor electrically coupled at its one end to the plurality ofelectromagnetic shields and wherein the other end of the groundconductor is configured to be electrically coupled to a housing of theactive implantable medical device.
 9. The implantable lead of claim 1,wherein the first inductive coil comprises at least two inductive coils,where each of the two inductive coils of the first inductive coil areself-resonant at or near a different MRI RE pulsed frequency.
 10. Theimplantable lead of claim 9, wherein one of the at least two inductivecoils of the first inductive coil is self-resonant at or near 128 MHzand the other of the at least two inductive coils of the first inductivecoil is self-resonant at or near 64 MHz.
 11. The implantable lead ofclaim 10, wherein the second inductive coil comprises at least twoinductive coils, where each of the two inductive coils of the secondinductive coil are self-resonant at or near a different MRI RF pulsedfrequency.
 12. The implantable lead of claim 11, wherein one of the atleast two inductive coils of the second inductive coil is self-resonantat or near 128 MHz and the other of the at least two inductive coils ofthe second inductive coil is self-resonant at or near 64 MHz.
 13. Theimplantable lead of claim 1, wherein the first inductive coil comprisesa plurality inductive coils and wherein the first electromagnetic shieldcomprises a plurality of electromagnetic shields electrically coupledtogether, where the plurality of electromagnetic shields are at leastpartially surrounding respectively the plurality of inductive coils. 14.The implantable lead of claim 1, wherein the first inductive coilcomprises 2, 3 or n number inductive coils and wherein the firstelectromagnetic shield comprises 2, 3 or n number of electromagneticshields electrically coupled together, where the 2, 3 or n number ofelectromagnetic shields are at least partially surrounding respectivelythe 2, 3 or n number inductive coils.
 15. The implantable lead of claim1, wherein the first inductive coil comprises a multi-layer coil. 16.The implantable lead of claim 1, including a second electromagneticshield at least partially surrounding the second inductive coil.
 17. Theimplantable lead of claim 16, wherein the first and secondelectromagnetic shields are electrically coupled together.
 18. Theimplantable lead of claim 17, including a ground conductor electricallycoupled at its one end to the first or second electromagnetic shield andwherein the other end of the ground conductor is configured to beelectrically coupled to a housing of the active implantable medicaldevice.
 19. The implantable lead of claim 1, wherein the tip electrodecomprises a translational active fixation helix tip electrode, andincluding a torque carrying support structure mechanically coupled tothe helix tip electrode and the first conductor, where the torquecarrying support structure is disposed mechanically in parallel with thefirst inductive coil.
 20. The implantable lead of claim 19, wherein thefirst inductive coil substantially surrounds a portion of the torquecarrying support structure.
 21. The implantable lead of claim 19,wherein the torque carrying support structure comprises a dielectricmaterial, an insulative material, a non-ferromagnetic material, aceramic material or a plastic material.
 22. The implantable lead ofclaim 21, wherein the torque carrying support structure is configured toreduce a torque experienced by the first inductive coil when the helixtip electrode is inserted into biological tissue.
 23. The implantablelead of claim 22, wherein the dielectric material of the torque carryingsupport structure comprises a dielectric constant greater than 1 up to100.
 24. The implantable lead of claim 23, wherein the dielectricmaterial of the support structure comprises a thermal-setting liquid.25. The implantable lead of claim 1, including a diverter elementelectrically coupled at its one end to the first conductor and the otherend of the diverter element is coupled to the first electromagneticshield.
 26. The implantable lead of claim 25, wherein the diverterelement comprises a capacitor, a capacitance, a parasitic capacitance oran L-C trap filter.
 27. The implantable lead of claim 26, wherein thediverter element is at least partially disposed within the firstelectromagnetic shield.
 28. The implantable lead of claim 1, including adiverter element electrically coupled at its one end to the firstconductor and the other end of the diverter element is coupled to thesecond conductor.
 29. The implantable lead of claim 28, wherein thediverter element comprises a capacitor, a capacitance, a parasiticcapacitance or an L-C trap filter.
 30. The implantable lead of claim 29,wherein the diverter element is at least partially disposed within thefirst electromagnetic shield.
 31. The implantable lead of claim 1,wherein the first electromagnetic shield is selected from the groupconsisting of a solid conductor, a wound spiral conductor, a mesh, atubing, a nanocoating, a conductive heat-shrink tubing, a conductivefoil, a wire, a braid, a circuit trace, a solid tubular material, aconductive polymer, a conductive epoxy, a carbon nano-fiber, anano-mesh, a nano-coating, and a nano-thread.
 32. The implantable leadof claim 1, wherein the first electromagnetic shield comprises abiocompatible metal or alloy selected from the group consisting oftitanium, platinum, platinum-iridium, gold, palladium, tantalum, carbon,niobium and alloys thereof.
 33. The implantable lead of claim 1, whereinthe first electromagnetic shield comprises a non-metallic materialselected from the group consisting of sapphire, ruby, lumina, ceramicand glass, and wherein the non-metallic material includes a thin layerof conductive metal deposited on either an inside or an outside surface.